Typically, in computed tomography (CT) imaging systems, a rotatable gantry includes an x-ray tube, detector, data acquisition system (DAS), and other components that rotate about a patient that is positioned at the approximate rotational center of the gantry. X-rays emit from the x-ray tube, are attenuated by the patient, and are received at the detector. The detector typically includes a photodiode-scintillator array of pixelated elements that convert the attenuated x-rays into photons within the scintillator, and then to electrical signals within the photodiode. The electrical signals are digitized and then received within the DAS, processed, and the processed signals are transmitted via a slipring (from the rotational side to the stationary side) to a computer or data processor for image reconstruction, where an image is formed.
The gantry typically includes a pre-patient collimator that defines or shapes the x-ray beam emitted from the x-ray tube. X-rays passing through the patient can cause x-ray scatter to occur, which can cause image artifacts. Thus, x-ray detectors typically include an anti-scatter grid (ASG) for collimating x-rays received at the detector. Imaging data may be obtained using x-rays that are generated at a single polychromatic energy. However, some systems may obtain multi-energy images that provide additional information for generating images.
Third generation multi-slices CT scanners typically include a detector assembly having scintillator/photodiodes arrays positioned in an arc, where the focal spot is the center of the corresponding circle. The material used in these detectors generally has scintillation crystal/photodiode arrays, where the scintillation crystal absorbs x-rays and converts the absorbed energy into visible light. A photodiode is used to convert the light to an electric current. The reading is typically proportional and linear to the total energy absorbed in the scintillator.
To ensure good image quality, third generation CT scanners, and particularly detector modules within the detector assembly, should satisfy very strict specifications that include but are not necessarily limited to: a) stability of the detector over time and temperature; b) in-sensitivity to focal spot motion; and c) stable and high light output over lifetime of the detector; as examples. As such, CT scanners are typically calibrated to account for the above items.
Nevertheless, the detectors can fall out of calibration due to radiation damage to the detectors, changes in operating conditions (such as temperature) of the CT scanner, and aging of the x-ray tube, as examples. To maintain proper calibration, some known CT systems therefore include a reference detector that may be used during acquisition of CT imaging data. For instance, some existing technologies use a detector located in the x-ray tube housing or at the pre-patient collimator, and others use a detector located outside the field-of-view FOV, such as at the edges of the detector, in order to avoid blockage from the patient.
However, although such detectors may provide a convenient and readily available reference for calibration purposes, such detectors are subject to accelerated aging and damage because, being outside the FOV, they typically experience increased radiation dose, and damage can result to the reference detector during the life of the detector assembly.
Thus, there is a need to improve detector calibration in a CT system.